Device and method for determining analyte levels

ABSTRACT

Devices and methods for determining analyte levels are described. The devices and methods allow for the implantation of analyte-monitoring devices, such as glucose monitoring devices that result in the delivery of a dependable flow of blood to deliver sample to the implanted device. The devices include unique architectural arrangement in the sensor region that allows accurate data to be obtained over long periods of time.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of Ser. No. 12/696,003, filed Jan.28, 2010, which is a continuation of Ser. No. 11/546,157, filed Oct. 10,2006, which is a continuation of Ser. No. 11/039,269, filed Jan. 19,2005, now U.S. Pat. No. 7,136,689, which is a continuation of Ser. No.09/916,858, filed Jul. 27, 2001, now U.S. Pat. No. 6,862,465, which is acontinuation-in-part of Ser. No. 09/447,227, filed Nov. 22, 1999, whichis a divisional of Ser. No. 08/811,473, filed Mar. 4, 1997, now U.S.Pat. No. 6,001,067.

FIELD OF THE INVENTION

The present invention relates generally to devices and methods fordetermining analyte levels, and, more particularly, to implantabledevices and methods for monitoring glucose levels in a biological fluid.

BACKGROUND OF THE INVENTION

The continuous measurement of substances in biological fluids is ofinterest in the control and study of metabolic disorders. Electrodesystems have been developed for this purpose whereby an enzyme-catalyzedreaction is monitored (e.g., by the changing concentrations of reactantsor products) by an electrochemical sensor. In such electrode systems,the electrochemical sensor comprises an electrode with potentiometric oramperometric function in close contact with a thin layer containing anenzyme in dissolved or insoluble form. Generally, a semipermeablemembrane separates the thin layer of the electrode containing the enzymefrom the sample of biological fluid that includes the substance to bemeasured.

Electrode systems that include enzymes have been used to convertamperometrically inactive substances into reaction products that areamperometrically active. For example, in the analysis of blood forglucose content, glucose (which is relatively inactive amperometrically)may be catalytically converted by the enzyme glucose oxidase in thepresence of oxygen and water to gluconic acid and hydrogen peroxide.Tracking the concentration of glucose is thereby possible since forevery glucose molecule reacted a proportional change in either oxygen orhydrogen peroxide sensor current will occur [U.S. Pat. Nos. 4,757,022and 4,994,167 to Shults et al., both of which are hereby incorporated byreference]. Hydrogen peroxide is anodically active and produces acurrent that is proportional to the concentration of hydrogen peroxide.[Updike et al., Diabetes Care, 11:801-807 (1988)].

Despite recent advances in the field of implantable glucose monitoringdevices, presently used devices are unable to provide data safely andreliably for long periods of time (e.g., months or years) [See, e.g.,Moatti-Sirat et al., Diabetologia 35:224-30 (1992)]. For example, Armouret al., Diabetes 39:1519-26 (1990), describes a miniaturized sensor thatis placed intravascularly, thereby allowing the tip of the sensor to bein continuous contact with the blood. Unfortunately, probes that areplaced directly into the vasculature put the recipient at risk forthrombophlebosis, thromboembolism, and thrombophlebitis.

Currently available glucose monitoring devices that may be implanted intissue (e.g., subcutaneously) are also associated with severalshortcomings. For example, there is no dependable flow of blood todeliver sample to the tip of the probe of the implanted device.Similarly, in order to be effective, the probe must consume some oxygenand glucose, but not enough to perturb the available glucose which it isintended to measure; subcutaneously implanted probes often reside in arelatively low oxygen environment in which oxygen or glucose depletionzones around the probe tip may result in erroneously low measuredglucose levels. In addition, implantable devices that utilize electrodesensors require membranes of the appropriate composition to protect thesensor from harsh in vivo environmental conditions. Current membranetechnology has allowed the development of a single structural membranethat performs the same functions that previously required multiplemembranes. However, these single membranes have been observed todelaminate and thus prevent accurate long-term glucose monitoring.Finally, the probe may be subject to “motion artifact” because thedevice is not adequately secured to the tissue, thus contributing tounreliable results. Partly because of these limitations, it haspreviously been difficult to obtain accurate information regarding thechanges in the amounts of analytes (e.g., whether blood glucose levelsare increasing or decreasing); this information is often extremelyimportant, for example, in ascertaining whether immediate correctiveaction is needed in the treatment of diabetic patients.

There is a need for a device that accurately and continuously determinesthe presence and the amounts of a particular analyte, such as glucose,in biological fluids. The device should be easy to use, be capable ofaccurate measurement of the analyte over long periods of time, andshould not readily be susceptible to motion artifact.

SUMMARY OF THE INVENTION

The present invention relates generally to devices and methods fordetermining analyte levels, and, more particularly, to implantabledevices and methods for monitoring glucose levels in a biological fluid.

In one aspect of the present invention, an implantable device formeasuring an analyte in a biological fluid is provided, which includesthe following: a housing including an electronic circuit; and a sensoroperably connected to the electronic circuit of the housing, the sensorincluding i) a member for determining the amount of glucose in abiological sample ii) a bioprotective membrane, the bioprotectivemembrane positioned more distal to the housing than the glucosedetermining member and substantially impermeable to macrophages, andiii) an angiogenic layer, the angiogenic layer positioned more distal tothe housing than the bioprotective membrane.

The present invention further encompasses a method of monitoring glucoselevels, the method including the steps of providing a host, and animplantable device as described above and implanting the device in thehost under conditions such that the device measures glucose for a periodexceeding 360 days.

In one embodiment of this aspect, the invention encompasses a method ofmeasuring glucose in a biological fluid that includes the steps ofproviding a host, and an implantable device as provided above, whereinthe glucose determining member of the implantable device is capable ofcontinuous glucose sensing, and implanting the device in the host.

DEFINITIONS

In order to facilitate an understanding of the present invention, anumber of terms are defined below.

The term “accurately” means, for example, 95% of measured values within25% of the actual value as determined by analysis of blood plasma,preferably within 15% of the actual value, and most preferably within 5%of the actual value. Alternatively, “accurately” means that 85% of themeasured values fall into the A and B regions of a Clarke error grid, orpreferably 90%, or most preferably 95% of the measured values fall intothese regions. It is understood that like any analytical device,calibration, calibration validation and recalibration are required forthe most accurate operation of the device.

The term “analyte” refers to a substance or chemical constituent in abiological fluid (e.g., blood or urine) that can be analyzed. Apreferred analyte for measurement by the devices and methods of thepresent invention is glucose.

The terms “sensor interface,” “sensor means,” “sensor” and the likerefer to the region of a monitoring device responsible for the detectionof a particular analyte. For example, in some embodiments of a glucosemonitoring device, the sensor interface refers to that region wherein abiological sample (e.g., blood or interstitial fluid) or a portionthereof contacts (directly or after passage through one or moremembranes or layers) an enzyme (e.g., glucose oxidase); the reaction ofthe biological sample (or portion thereof) results in the formation ofreaction products that allow a determination of the glucose level in thebiological sample. In preferred embodiments of the present invention,the sensor means comprises an angiogenic layer, a bioprotective layer,an enzyme layer, and an electrolyte phase (i.e., a free-flowing liquidphase comprising an electrolyte-containing fluid [described furtherbelow]). In some preferred embodiments, the sensor interface protrudesbeyond the plane of the housing.

The term “tissue interface” refers to that region of a monitoring devicethat is in contact with tissue.

The terms “operably connected,” “operably linked,” and the like refer toone or more components being linked to another component(s) in a mannerthat allows transmission of, e.g., signals between the components. Forexample, one or more electrodes may be used to detect the amount ofanalyte in a sample and convert that information into a signal; thesignal may then be transmitted to electronic circuit means (i.e., theelectrode is “operably linked” to the electronic circuit means), whichmay convert the signal into a numerical value in the form of knownstandard values.

The term “electronic circuit means” or “electronic circuit” refers tothe electronic circuitry components of a biological fluid measuringdevice required to process information obtained by a sensor meansregarding a particular analyte in a biological fluid, thereby providingdata regarding the amount of that analyte in the fluid. U.S. Pat. No.4,757,022 to Shults et al., previously incorporated by reference,describes suitable electronic circuit means (see, e.g., FIG. 7); ofcourse, the present invention is not limited to use with the electroniccircuit means described therein. A variety of circuits are contemplated,including but not limited to those circuits described in U.S. Pat. Nos.5,497,772 and 4,787,398, hereby incorporated by reference.

The terms “angiogenic layer,” “angiogenic membrane,” and the like referto a region, membrane, etc. of a biological fluid measuring device thatpromotes and maintains the development of blood vessels microcirculationaround the sensor region of the device. As described in detail below,the angiogenic layer of the devices of the present invention may beconstructed of membrane materials alone or in combination such aspolytetrafluoroethylene, hydrophilic polyvinylidene fluoride, mixedcellulose esters, polyvinylchloride, and other polymers including, butnot limited to, polypropylene, polysulfone, and polymethylmethacrylate.

The phrase “positioned more distal” refers to the spatial relationshipbetween various elements in comparison to a particular point ofreference. For example, some embodiments of a biological fluid measuringdevice comprise both a bioprotective membrane and an angiogeniclayer/membrane. If the housing of the biological fluid measuring deviceis deemed to be the point of reference and the angiogenic layer ispositioned more distal to the housing than the bioprotective layer, thenthe bioprotective layer is closer to the housing than the angiogeniclayer.

The terms “bioprotective membrane,” “bioprotective layer,” and the likerefer to a semipermeable membrane comprised of protective biomaterialsof a few microns thickness or more that are permeable to oxygen andglucose and are placed over the tip of the sensor to keep the whiteblood cells (e.g., tissue macrophages) from gaining proximity to andthen damaging the enzyme membrane. In some embodiments, thebioprotective membrane has pores (typically from approximately 0.1 toapproximately 1.0 micron). In preferred embodiments, a bioprotectivemembrane comprises polytetrafluoroethylene and contains pores ofapproximately 0.4 microns in diameter. Pore size is defined as the poresize provided by the manufacturer or supplier.

The phrase “substantially impermeable to macrophages” means that few, ifany, macrophages are able to cross a barrier (e.g., the bioprotectivemembrane). In preferred embodiments, fewer than 1% of the macrophagesthat come in contact with the bioprotective membrane are able to cross.

The phrase “material for securing said device to biological tissue”refers to materials suitable for attaching the devices of the presentinvention to, the fibrous tissue of a foreign body capsule. Suitablematerials include, but are not limited to, poly(ethylene terephthalate).In preferred embodiments, the top of the housing is covered with thematerials in the form of surgical grade fabrics; more preferredembodiments also contain material in the sensor interface region (seeFIG. 1B).

The phrase “member for determining the amount of glucose in a biologicalsample” refers broadly to any mechanism (e.g., enzymatic ornon-enzymatic) by which glucose can be quantitated. For example, someembodiments of the present invention utilize a membrane that containsglucose oxidase that catalyzes the conversion of glucose to gluconate:Glucose+O₂=Gluconate+H₂O₂. Because for each glucose molecule convertedto gluconate, there is a proportional change in the co-reactant O₂ andthe product H₂O₂, one can monitor the current change in either theco-reactant or the product to determine glucose concentration.

The phrase “apparatus for transmitting data to a location external tosaid device” refers broadly to any mechanism by which data collected bya biological fluid measuring device implanted within a subject may betransferred to a location external to the subject. In preferredembodiments of the present invention, radiotelemetry is used to providedata regarding blood glucose levels, trends, and the like.

The terms “radiotelemetry,” “radiotelemetric device,” and the like referto the transmission by radio waves of the data recorded by the implanteddevice to an ex vivo recording station (e.g., a computer), where thedata is recorded and, if desired, further processed (see, e.g., U.S.Pat. Nos. 5,321,414 and 4,823,808, hereby incorporated by reference; PCTPat. Publication WO 94/22367).

The term “host” refers to both humans and animals.

The phrase “continuous glucose sensing” refers to the period in whichmonitoring of plasma glucose concentration is continuously carried out.More specifically, at the beginning of the period in which continuousglucose sensing is effected, the background sensor output noisediminishes and the sensor output stabilizes (e.g., over several days) toa long-term level reflecting adequate microcirculatory delivery ofglucose and oxygen to the tip of the sensor (see FIG. 2).

The term “filtrate layer” refers to any permeable membrane that is ableto limit molecules from passing through the membrane based on theirproperties including molecular weight. More particularly, the resistancelayer, interference layer and bioprotective membrane are examples oflayers that can function as filtrate layers, depending on the materialsfrom which they are prepared. These layers can control delivery ofanalyte to a sensing means. Furthermore, these layers can reduce anumber of undesirable molecular species that may otherwise be exposed tothe sensor for detection and provide a controlled sample volume to theanalyte sensing means.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A depicts a cross-sectional drawing of one embodiment of animplantable analyte measuring device of the present invention.

FIG. 1B depicts a cross-sectional exploded view of the sensor interfacedome of FIG. 1A.

FIG. 1C depicts a cross-sectional exploded view of theelectrode-membrane region of FIG. 1B detailing the sensor tip and thefunctional membrane layers.

FIG. 2 graphically depicts glucose levels as a function of the number ofdays post-implant.

FIG. 3 is a graphical representation of the number of functional sensorsversus time (i.e. weeks) comparing control devices including only acell-impermeable domain (“Control”), with devices including acell-impermeable domain and a barrier-cell domain (“Test”).

FIG. 4A is a photograph of an intact composite bioprotective/angiogenicmembrane after implantation in a dog for 137 days.

FIG. 4B is a photograph of a delaminated ePTFE bilayer membrane afterimplantation in a dog for 125 days.

DETAILED DESCRIPTION OF THE INVENTION

The present invention relates generally to devices and methods fordetermining analyte levels, and, more particularly, to implantabledevices and methods for monitoring glucose levels in a biological fluid.In a preferred embodiment, the device and methods of the presentinvention are used to determine the level of glucose in a host, aparticularly important measurement for individuals having diabetes.

Although the description that follows is primarily directed at glucosemonitoring devices and methods for their use, the devices and methods ofthe present invention are not limited to glucose measurement. Rather,the devices and methods may be applied to detect and quantitate otheranalytes present in biological fluids (including, but not limited to,amino acids and lactate), especially those analytes that are substratesfor oxidase enzymes [see, e.g., U.S. Pat. No. 4,703,756 to Gough et al.,hereby incorporated by reference]. Moreover, the devices and methods ofthe present invention may be utilized to present components ofbiological fluids to measurement methods which are not enzyme-based,including, but not limited to, those based on surface plasmon resonance,surface acoustic waves, optical absorbance in the long wave infraredregion, and optical rotation of polarized light.

For example, surface plasmon resonance sensors that analyze a regionwithin less than one wavelength of analysis light near the flat surfaceof the sensor have been described (See U.S. Pat. No. 5,492,840). Thesesensors have been used, for example, in the study of immunochemistry andother surface bound chemical reactions (Jonsson et al., Annales deBiologies Clinique 51 (10:19, 1993). This type of sensor may beincorporated into the implantable device of the present invention forthe detection of a number of different analytes including glucose. Oneskilled in the art would recognize that the surface plasmon resonancesensor is an optical sensor and that the implantable device of thepresent invention may further include a source of coherent radiation(e.g. a laser operating in the visible or near infrared).

In one Application, referred to here as a consumptive approach, anenzyme that consumes the analyte producing a detectable product isimmobilized on the sensor in the filtrate layer. When the enzymeconsumes the analyte, the reaction products diffuse away from the enzymeat a rate dependent on the permeability of the layers distal to theenzyme layer. As a result, reaction products will accumulate at a higherconcentration near the sensor, within one wavelength of analysis light,where they may be detected and measured. One example of such a systemthat detects the presence of glucose would immobilize a glucose oxidaseenzyme layer on the sensor surface.

The layers of the present invention play an important role in theeffective operation and function of this type of sensor. In particular,the angiogenic layer assures a constant supply of analyte from thetissues of the subject, the bioprotective membrane protects theunderlying layers from cellular attack, the resistance layer controlsthe rate of delivery of analyte and the filtrate layer performs manyfunctions including; providing a low molecular weight filtrate, reducingthe number of undesirable molecular species available to the sensor fordetection and providing a controlled volume of sample for detection bythe sensor. As mentioned above, the bioprotective membrane, resistancelayer and interference layer can function as filtrate layers. Forexample, it is well within the contemplation of the present inventionthat the bioprotective membrane can be made of a material that is ableto exclude certain molecules from passing through the membrane based ontheir size.

One skilled in the art would recognize that the reaction kineticsassociated with each type of enzyme that may be selected for use withthis sensor is unique. However, in general, if an excess of enzyme isprovided, the enzyme turnover rate is proportional to the flux ofanalyte to the enzyme and independent of the enzyme concentration.Therefore, the actual analyte concentration may be calculated utilizingthe diffusion rate of the detectable analyte across the bioprotectiveresistance layers.

In another application, referred to here as a non-consumptive approach,an analyte-binding compound is provided on the surface plasmon resonancesensor surface within one wavelength of analysis light. This compoundreversibly binds, but does not consume, the analyte. In thisapplication, the analyte moves reversibly onto and off of attachmentsites on the binding compound. This reaction provides a steady statecondition for bound and unbound analyte that may be quantitated andanalyte concentration mathematically calculated. One skilled in the artwould recognize that the reaction kinetics associated with binding andrelease of the analyte is unique for each type of binding compoundselected. Examples of such a system that detects the presence of glucoseprovide a binding compound comprised of conconavalin A or a wide rangeof borate containing compounds (See U.S. Pat. No. 6,011,985).

Since this is a chemical equilibrium-based approach, a filtrate layer isnot necessarily required to maintain an analyte concentration near thesensor. However, such a membrane would still be desired to reduce thenumber of undesirable molecular species available to the analyte-bindinglayer. Preferably, the bioprotective layer is thin to allow rapid sensorequilibration to changes in analyte levels. As described above, oneskilled in the art would recognize that the function of the filtratelayer could be incorporated into the bioprotective membrane by selectionof the appropriate molecular exclusion, such as exclusion by molecularweight, if desired.

A variety of materials may be utilized to construct a combinationangiogenic/bioprotective membrane, many of which are described belowunder the angiogenic layer and bioprotective membrane headings.Preferably, this combination membrane is ePTFE embedded in a layer ofPVP containing urethane hydrogel. However, any material that performs asimilar function as the PVP containing polyurethane hydrogel could besubstituted.

In either application, consumptive or non-consumptive, one skilled inthe art would recognize that the response time of the sensor is subjectto Fick's law of diffusion. More specifically, sensors with thickmembrane layers or that have low analyte diffusivity will respond slowerto change in analyte concentration than sensors with thin membranes orthat have high analyte diffusivity. Consequently, reasonableoptimization experimentation with the membrane and layers would berequired to meet various use requirements.

One skilled in the art would further recognize that the consumptive ornon-consumptive approaches of the previous example could be applied toadditional sensor modalities as follows:

1. Another sensor that may be incorporated into the device of thepresent invention that has been previously described is a surfaceacoustic wave sensor (See U.S. Pat. No. 5,932,953). This sensor, alsoreferred to as a bulk-acoustic wave piezoelectric resonator, typicallyincludes a planar surface of piezoelectric material with two respectivemetal layers bonded on opposite sides that form the electrodes of theresonator. The two surfaces of the resonator are free to undergovibrational movement when the resonator is driven by a signal within theresonance band of the resonator. One of these surfaces is adapted toprovide reversible binding sites for the analyte being detected. Thebinding of the analyte on the surface of the resonator alters theresonant characteristics of the resonator and changes in the resonantcharacteristics may be detected and interpreted to provide quantitativeinformation regarding the analyte concentration.

2. Another sensor that may be incorporated into the device of thepresent invention is an optical absorbance sensor (See U.S. Pat. No.6,049,727). This sensor utilizes short to medium wavelength infraredlight that is passed through a sample with the unabsorbed infrared lightbeing monitored by an optical detector.

Previously developed methods for analysis of analytes such as glucose intissues and blood have been relatively unsuccessful for two reasons,interference from other chemicals present in the complex biologicalsample and signal variation due to poor control of sample volume. Theseproblems may be solved by providing a low molecular weight filtrate ofbiological fluid in a controlled volume of sample to the sensor. In onesystem of the present invention, biological analyte is provided to thesensor through the angiogenic layer. This analyte is then filteredthrough the bioprotective membrane to produce a desirable filtrate.Alternatively, a third filtrate layer, such as an interference layer,may be utilized having specific filtration properties to produce thedesired filtrate. The three-dimensional structure of the bioprotectivemembrane and/or other filtrate layers is utilized to define andstabilize the sample volume. One skilled in the art would recognize thatany material that provides a low molecular weight filtrate to the sensorin a controlled volume might be utilized. Preferably, this material ispolyurethane.

The sensor may be enhanced by partial metallization of the distal sideof the filtrate producing material that would serve to isolate byreflection the optical signal to the space within the filtrate regiondirectly adjacent to the sensor. This metal film may be a durable metalincluding, but not restricted to, gold or platinum and may be vacuumdeposited onto the filtrate producing material.

One skilled in the art would recognize that the optical absorbancesensor requires a source of short to medium wavelength infrared light.Consequently, the implantable device of the present invention wouldfurther include a source of infrared radiation and an optical detector.

3. Another sensor that may be incorporated into the device of thepresent invention that has been previously described is a polarizedlight optical rotation sensor (See U.S. Pat. No. 5,209,231). This sensormay be used to detect an analyte that rotates polarized light such asglucose. In particular, glucose concentrations in biological fluids inthe range of 0.05 to 1.00% w/v may be detected and quantitated. Normalnon-diabetic subjects generally have biological glucose concentrationsranging from 0.07 to 0.12% w/v.

In this type of sensor, the optical detector receives polarized lightpassed through a biological sample and then further through a polarizingfilter. The optical activity of an analyte in the sample rotates thepolarized light in proportion to its concentration. Unfortunately,accurate measurements of glucose in complex biological samples hasproven difficult because of the optical activity of interferingsubstances and poor control of sample volume. These problems may besolved by providing a low molecular weight filtrate of biological fluidin a controlled volume to the sensor. The present invention meets thiscriterion by providing a continuous supply of biological glucose to thesensor through the angiogenic layer that is filtered through abioprotective membrane and/or a filtrate layer as described previouslyfor the optical absorbance sensor. One skilled in the art wouldrecognize that any material that provides a low molecular weightfiltrate to the sensor in a controlled geometry might be utilized.Preferably, this material is polyurethane. In addition, one skilled inthe art would recognize that the polarized light optical rotation sensorrequires a source of polarized light. Consequently, the implantabledevice of the present invention would further include a source ofpolarized radiation.

4. Another sensor that may be incorporated into the device of thepresent invention that has been previously described is a fluorescencesensor (See U.S. Pat. No. 5,341,805). The invention of Colvin provides amethod for incorporating an ultraviolet light source and fluorescentsensing molecules in an implantable device. However, Colvin does notdescribe how the sensor would survive harsh in vivo environmentalconditions, how the device would be functionally integrated into bodytissues or how a continuous supply of glucose would be maintained fordetection by the sensor. These problems may be solved by providing a lowmolecular weight filtrate of biological fluid in a controlled volume tothe sensor.

In this example, a continuous supply of biological glucose passes to thesensor through the angiogenic layer that prevents isolation of thesensor by the body tissue. The glucose is then filtered through thebioprotective membrane to produce a desirable filtrate having fewerinterfering molecules and to protect the sensor from in vivoenvironmental conditions. Alternatively, a filtrate layer may beutilized having specific filtration properties to produce the desiredfiltrate. The three-dimensional structure of the bioprotective membraneand/or filtrate layer also provides stabilized sample volume fordetection by the sensor.

One skilled in the art would recognize that a fluorescence sensorrequires a source of light. Consequently, the implantable device of thepresent invention would further comprise a source of radiation, as wellas fluorescent sensing molecules to detect the presence of analyte.

I. NATURE OF THE FOREIGN BODY CAPSULE

Devices and probes that are implanted into subcutaneous tissue willalmost always elicit a foreign body capsule (FBC) as part of the body'sresponse to the introduction of a foreign material. Therefore,implantation of a glucose sensor results in an acute inflammatoryreaction followed by building of fibrotic tissue. Ultimately, a matureFBC including primarily a vascular fibrous tissue forms around thedevice (Shanker and Greisler, Inflammation and Biomaterials in Greco RS, ed. Implantation Biology: The Host Response and Biomedical Devices,pp 68-80, CRC Press (1994)).

Although fluid is frequently found within the capsular space between thesensor and the capsule, levels of analytes (e.g., glucose and oxygen)within the fluid often do not mimic levels in the body's vasculature,making accurate measurement difficult.

In general, the formation of a FBC has precluded the collection ofreliable, continuous information, reportedly because of poorvascularization, the composition of a FBC has prevented stabilization ofthe implanted device, contributing to motion artifact that rendersunreliable results. Thus, conventionally, it has been the practice ofthose skilled in the art to attempt to minimize FBC formation by, forexample, using a short-lived needle geometry or sensor coatings tominimize the foreign body reaction (“Biosensors in the Body” David M.Fraser, ed.; 1997 pp 117-170. Wiley & Sons Ltd., West Sussex, England).

In contrast to the prior art, the teachings of the present inventionrecognize that FBC formation is the dominant event surrounding long termimplantation of any sensor and must be orchestrated to support ratherthan hinder or block sensor performance. For example, sensors often donot perform well until the FBC has matured sufficiently to provideingrowth of well-attached tissue bearing a rich supply of capillariesdirectly to the surface of the sensor. With reference to FIG. 2,stabilization of device function generally occurs between about 2 and 8weeks depending on the rate of healing and formation of new capillaries.In some cases, devices are functional from the time of implant, andsometimes it may take as long as 12 weeks. However, the majority ofdevices begin functioning between weeks 2 and 8 after implantation. Thismaturation process, when initiated according to the present invention,is a function of biomaterial and host factors that initiate and modulateangiogenesis, and promote and control fibrocyte ingrowth. The presentinvention contemplates the use of particular materials to promoteangiogenesis adjacent to the sensor interface (also referred to as theelectrode-membrane region, described below) and to anchor the devicewithin the FBC.

II. THE IMPLANTABLE GLUCOSE MONITORING DEVICE OF THE PRESENT INVENTION

The present invention contemplates the use of a unique micro-geometry atthe sensor interface of an implantable device. Moreover, the presentinvention contemplates the use of materials covering all or a portion ofthe device to assist in the stabilization of the device followingimplantation. However, it should be pointed out that the presentinvention does not require a device comprising particular electroniccomponents (e.g., electrodes, circuitry, etc). Indeed, the teachings ofthe present invention can be used with virtually any monitoring devicesuitable for implantation (or subject to modification allowingimplantation); suitable devices include, but are not limited, to thosedescribed in U.S. Pat. No. 6,001,067 to Shults et al.; U.S. Pat. No.4,703,756 to Gough et al., and U.S. Pat. No. 4,431,004 to Bessman etal.; the contents of each being hereby incorporated by reference, andBindra et al., Anal. Chem. 63:1692-96 (1991).

In the discussion that follows, an example of an implantable device thatincludes the features of the present invention is first described.Thereafter, the specific characteristics of, for example, the sensorinterface contemplated by the present invention will be described indetail.

Generally speaking, the implantable devices contemplated for use withthe present invention are cylindrical or oval shaped; of course, deviceswith other shapes may also be used with the present invention. Thesample device includes a housing composed of radiotransparent ceramic.FIG. 1A depicts a cross-sectional drawing of one embodiment of animplantable measuring device. Referring to FIG. 1A, the cylindricaldevice includes a ceramic body 1 and ceramic head 10 houses the sensorelectronics that include a circuit board 2, a microprocessor 3, abattery 4, and an antenna 5. Furthermore, the ceramic body 1 and head 10possess a matching taper joint 6 that is sealed with epoxy. Theelectrodes are subsequently connected to the circuit board via a socket8.

As indicated in detail in FIG. 1B, three electrodes protrude through theceramic head 10, a platinum working electrode 21, a platinum counterelectrode 22 and a silver/silver chloride reference electrode 20. Eachof these is hermetically brazed 26 to the ceramic head 10 and furtheraffixed with epoxy 28. The sensing region 24 is covered with the sensingmembrane described below and the ceramic head 10 contains a groove 29 sothat the membrane may be affixed into place with an o-ring.

In a preferred embodiment, the device is cylindrical, as shown in FIG.1A, and is approximately 1 cm in diameter, and 5.5 cm long. The sensingregion is situated at one extreme end of the cylinder. The sensor regionincludes a dome onto which the sensing membranes are attached.

Maintaining the blood supply near an implanted foreign body like animplanted analyte-monitoring sensor requires stable fixation of FBCtissue on the surface of the foreign body. This can be achieved, forexample, by using capsular attachment (anchoring) materials (e.g., thosematerials that includes the sensor interface and tissue anchoringlayers) developed to repair or reinforce tissues, including, but notlimited to, polyester (DACRON®; DuPont; poly(ethylene terephthalate))velour, expanded polytetrafluoroethylene (TEFLON/®; Gore),polytetrafluoroethylene felts, polypropylene cloth, and related porousimplant materials. In a preferred embodiment, porous silicone materialsare used for anchoring the device. In another embodiment, non-wovenpolyester fibers are used for anchoring the device. Tissue tends toaggressively grow into the materials disclosed above and form a strongmechanical bond (i.e., tissue anchoring); this fixation of the implantin its capsule is essential to prevent motion artifact or disturbance ofthe newly developed capillary blood supply.

In a preferred embodiment, the anchoring material is attached directlyto the body of the device. In the case of non-woven polyester fibers,they may be sutured into place by rolling the material onto thecircumferential periphery of the device and further encircling themembrane with PTFE sutures and tying the sutures to hold the membrane inplace. In another preferred embodiment, porous silicone is attached tothe surface of the cylindrical device using medical grade siliconeadhesive. In either case, the material may be further held in place byan o-ring (FIG. 1B).

As shown in FIG. 1A, the interior of the housing contains one or morebatteries 4 operably connected to an electronic circuit means (e.g., acircuit board 2), which, in turn, is operably connected to at least oneelectrode (described below); in another embodiment, at least twoelectrodes are carried by the housing. In a preferred embodiment, threeelectrodes are used. Any electronic circuitry and batteries that renderreliable, continuous, long-term (e.g., months to years) results may beused in conjunction with the devices of the present invention.

The housing of the devices of the present invention preferably contain abiocompatible ceramic material. A preferred embodiment of the devicecontains a radiofrequency transmitter and antenna within the body of theceramic device. Ceramic materials are radiotransparent and, therefore,are preferred over metals that are radioopaque. Ceramic materials arepreferred over plastic materials (which may also be radiotransparent)because they are more effective than plastics at preventing waterpenetration. In one embodiment of the invention, the ceramic head andbody are connected at an approximately 0.9 cm long taper joint sealedwith epoxy. In other embodiments, the head and body may be attached bysealing with metals to produce a completely hermetic package.

FIG. 1C depicts a cross-sectional exploded view of theelectrode-membrane region 24 set forth in FIG. 1B detailing the sensortip and the functional membrane layers. As depicted in FIG. 1C, theelectrode-membrane region includes several different membrane layers,the compositions and functions of which are described in detail below.The top ends of the electrodes are in contact with the electrolyte phase30, a free-flowing fluid phase. The electrolyte phase is covered by thesensing membrane 32 that contains an enzyme, e.g., glucose oxidase, andseveral functional polymer layers (as described below). In turn, acomposite bioprotective/angiogenic membrane 33 covers the sensingmembrane 32 and serves, in part, to protect the sensor from externalforces that may result in environmental stress cracking of the sensingmembrane 32.

In one preferred embodiment of the inventive device, each of themembrane layers is affixed to the ceramic head 10 in FIGS. 1A and 1B byan o-ring. The o-ring may be formed of fluoroelastomer.

The present invention contemplates the construction of the membranelayers of the sensor interface region using standard film coatingtechniques. This type of membrane fabrication facilitates control ofmembrane properties and membrane testing.

III. THE SENSOR INTERFACE REGION

As mentioned above and disclosed in FIG. 1C, in a preferred embodiment,the sensor interface region includes several different layers andmembranes that cover the electrodes of an implantable analyte-measuringdevice. The characteristics of these layers and membranes are nowdiscussed in more detail. The layers and membranes prevent directcontact of the biological fluid sample with the electrodes, whilepermitting selected substances (e.g., analytes) of the fluid to passtherethrough for electrochemical reaction with the electrodes.

Measurement of analyte in a filtrate of biological fluid samples hasbeen shown to be preferred over direct measurement of analyte inbiological fluid in order to minimize effects of interfering substancesand improve control of sample volume. It is well known in the art thatelectrode surfaces exposed to a wide range of biological molecules willsuffer poisoning of catalytic activity and failure. However, utilizingthe layers and membranes of the present invention, the activeelectrochemical surfaces of the sensor electrodes are preserved,allowing activity to be retained for extended periods of time in vivo.By limiting exposure of the platinum sensor surface to certain molecularspecies (e.g., molecules having a molecular weight below 34 Daltons, themolecular weight of hydrogen peroxide), in vivo sensor operating life inexcess of one year in canine subjects has been observed.

A. Angiogenic Layer

For implantable glucose monitoring devices, a sensor/tissue interfacemust be created which provides the sensor with oxygen and glucoseconcentrations comparable to that normally available to tissue comprisedof living cells. Absent such an interface, the sensor is associated withunstable and chaotic performance indicating that inadequate oxygenand/or glucose are reaching the sensor. The development of interfaces inother contexts has been reported. For example, investigators havedeveloped techniques that stimulate and maintain blood vessels inside aFBC to provide for the demanding oxygen needs of pancreatic isletswithin an implanted membrane. [See, e.g., Brauker et al., J. Biomed.Mater. Res. (1995) 29:1517-1524]. These techniques depend, in part, onthe use of a vascularizing layer on the exterior of the implantedmembrane. However, previously described implantable analyte-monitoringdevices have not been able to successfully maintain sufficient bloodflow to the sensor interface.

As described above, the outermost layer of the electrode-membrane regionincludes an angiogenic material. The angiogenic layer of the devices ofthe present invention may be constructed of membrane materials such ashydrophilic polyvinylidene fluoride (e.g., Durapore®; Millipore Bedford,Mass.), mixed cellulose esters (e.g., MF; Millipore Bedford, Mass.),polyvinyl chloride (e.g., PVC; Millipore Bedford, Mass.), and otherpolymers including, but not limited to, polypropylene, polysulphone, andpolymethylmethacrylate. Preferably, the thickness of the angiogeniclayer is about 10 μm to about 20 μm. The angiogenic layer comprisespores sizes of about 0.5 μm to about 20 μm, and more preferably about1.0 μm to about 10 μm, sizes that allow most substances to pass through,including, e.g., macrophages. The preferred material is expanded PTFE ofa thickness of about 15 μm and pore sizes of about 5 μm to about 10 μm.

To further promote stable foreign body capsule structure withoutinterfering with angiogenesis, an additional outermost layer of materialcomprised of a thin low-density non-woven polyester (e.g., manufacturedby Reemay) can be laminated over the preferred PTFE described above. Inpreferred embodiments, the thickness of this layer is about 120 μm. Thisadditional thin layer of material does not interfere with angiogenesisand enhances the manufacturability of the angiogenic layer. [See U.S.Pat. No. 5,741,330 to Brauker et al., hereby incorporated by reference;also U.S. Pat. No. 5,782,912, U.S. Pat. No. 5,800,529, U.S. Pat. No.5,882,354 U.S. Pat. No. 5,964,804 assigned to Baxter].

B. Bioprotective Membrane

The inflammatory response that initiates and sustains a FBC isassociated with both advantages and disadvantages. Some inflammatoryresponse is needed to create a new capillary bed in close proximity tothe surface of the sensor in order to i) continuously deliver adequateoxygen and glucose and ii) create sufficient tissue ingrowth to anchorthe implant and prevent motion artifact. On the other hand, inflammationis associated with invasion of tissue macrophages that have the abilityto biodegrade many artificial biomaterials (some of which were, untilrecently, considered nonbiodegradable). When activated by a foreignbody, tissue macrophages degranulate, releasing from their cytoplasmicmyeloperoxidase system hypochlorite (bleach), H₂O₂ and other oxidantspecies. Both hypochlorite and H₂O₂ are known to break down a variety ofpolymers, including polyurethane, by a phenomenon referred to asenvironmental stress cracking. [Phillips et al., J. Biomat. Appl.,3:202-227 (1988); Stokes, J. Biomat. Appl. 3:228-259 (1988)]. Indeed,environmental stress cracking has been shown to limit the lifetime andperformance of an enzyme-active polyurethane membrane stretched over thetip of a glucose sensor. [Updike et al., Am. Soc. Artificial InternalOrgans, 40:157-163 (1994)].

Because both hypochlorite and H₂O₂ are short-lived chemical species invivo, biodegradation will not occur if macrophages are kept a sufficientdistance from the enzyme active membrane. The present inventioncontemplates the use of a bioprotective membrane that allows transportof glucose and oxygen but prevents the entry of inflammatory cells suchas macrophages and foreign body giant cells. The bioprotective membraneis placed proximal to the angiogenic membrane. It may be simply placedadjacent to the angiogenic layer without adhering, or it may be attachedwith an adhesive material to the angiogenic layer, or it may be cast inplace upon the angiogenic layer as described in Example 1. The devicesof the present invention are not limited by the nature of thebioprotective layer. However, the bioprotective layer should bebiostable for long periods of time (e.g., several years); the presentinvention contemplates the use of polymers including, but not limitedto, polyurethane, polypropylene, polysulphone, polytetrafluoroethylene(PTFE), and poly(ethylene terephthalate) (PET).

The bioprotective membrane and the angiogenic layer may be combined intoa single bilayer membrane as more fully described in Example 1. Theactive angiogenic function of the combined membrane is based on thepresentation of the ePTFE side of the membrane to the reactive cells ofthe foreign body capsule and further to the response of the tissue tothe microstructure of the ePTFE. This bioprotective/angiogenic membraneis unique in that the membrane does not delaminate as has been observedwith other commercially available membranes (see FIG. 4A as comparedwith FIG. 4B). This is desirable for an implantable device to assureaccurate measurement of analyte over long periods of time. Although thephysical structure of the ePTFE represents a preferred embodiment, manyother combinations of materials that provide the same function as themembrane of Example 1 could be utilized. For example, the ePTFE could bereplaced by other fine fibrous materials. In particular, polymers suchas spun polyolefin or non-organic materials such as mineral or glassfibers may be useful. Likewise, the polyurethane bioprotective layer ofExample 1, which includes a biostable urethane and polyvinylpyrrolidone(PVP), could be replaced by polymers able to pass analyte while blockingmacrophages and mechanically retaining the fine fibrous materialpresented to the reactive cells of the foreign body capsule.

C. Sensing Membrane

The present invention contemplates membranes impregnated with enzyme. Itis not intended that the present invention be limited by the nature ofthe enzyme membrane. The sensing membrane of a preferred embodiment isdepicted in FIG. 1C as being a single, homogeneous structure. However,in preferred embodiments, the sensing membrane includes a plurality ofdistinct layers. In a particularly preferred embodiment, the sensingmembrane includes the following four layers (in succession from thebioprotective membrane to the layer most proximal to the electrodes): i)a resistance layer; ii) an enzyme layer; iii) an interference layer; andiv) an electrolyte layer.

Resistance Layer

There is a molar excess of glucose relative to the amount of oxygen insamples of blood. Indeed, for every free oxygen molecule inextracellular fluid, there are typically more than 100 glucose moleculespresent [Updike et al., Diabetes Care 5:207-21 (1982)]. However, animmobilized enzyme-based sensor using oxygen (O₂) as cofactor must besupplied with oxygen in non-rate-limiting excess in order to respondlinearly to changes in glucose concentration while not responding tochanges in oxygen tension. More specifically, when a glucose-monitoringreaction is oxygen-limited, linearity is not achieved above minimalconcentrations of glucose. Without a semipermeable membrane over theenzyme layer, linear response to glucose levels can be obtained only upto about 40 mg/dL; however, in a clinical setting, linear response toglucose levels are desirable up to at least about 500 mg/dL.

The resistance layer includes a semipermeable membrane that controls theflux of oxygen and glucose to the underlying enzyme layer (i.e., limitsthe flux of glucose), rendering the necessary supply of oxygen innon-rate-limiting excess. As a result, the upper limit of linearity ofglucose measurement is extended to a much higher value than that whichcould be achieved without the resistance layer. The devices of thepresent invention contemplate resistance layers comprising polymermembranes with oxygen-to-glucose permeability ratios of approximately200:1; as a result, one-dimensional reactant diffusion is adequate toprovide excess oxygen at all reasonable glucose and oxygenconcentrations found in the subcutaneous matrix [Rhodes et al., Anal.Chem., 66:1520-1529 (1994)].

In preferred embodiments, the resistance layer has a thickness of lessthan about 45 more preferably in the range of about 15 to about 40 andmost preferably in the range of about 20 to about 35 μm.

The resistance layer is desirably constructed of a mixture ofhydrophobic and hydrophilic polyurethanes.

Enzyme Layer

In addition to glucose oxidase, the present invention contemplates theuse of a membrane layer impregnated with other oxidases, e.g., galactoseoxidase, uricase. For an enzyme-based electrochemical glucose sensor toperform well, the sensor's response must, neither be limited by enzymeactivity nor cofactor concentration. Because enzymes, including the veryrobust glucose oxidase, are subject to deactivation as a function ofambient conditions, this behavior needs to be accounted for inconstructing sensors for long-term use.

Excess glucose oxidase loading is required for long sensor life. Whenexcess glucose oxidase is used, up to 1.5 years of performance may bepossible from the glucose-monitoring devices contemplated by the presentinvention.

In one preferred embodiment, the enzyme layer includes a polyurethanelatex.

Interference Layer

The interference layer includes a thin, hydrophobic membrane that isnon-swellable and restricts diffusion of low molecular weight species.The interference layer is permeable to relatively low molecular weightsubstances, such as hydrogen peroxide, but restricts the passage ofhigher molecular weight substances, including glucose and ascorbic acid.The interference layer serves to allow analytes and other substancesthat are to be measured by the electrodes to pass through, whilepreventing passage of other substances.

Preferred materials from which the interference layer can be madeinclude polyurethanes. In one desired embodiment, the interference layerincludes an aliphatic polyetherurethane.

The interference layer has a preferred thickness of less than about 5μm, more preferably in the range of about 0.1 to about 5 μm and mostpreferably in the range of about 0.5 to about 3 μm. Thicker membranesalso may be useful, but thinner membranes are preferred because theyhave a lower impact on the rate of diffusion of hydrogen peroxide fromthe enzyme membrane to the electrodes.

Electrolyte Layer

To ensure electrochemical reaction, the electrolyte layer comprises asemipermeable coating that maintains hydrophilicity at the electroderegion of the sensor interface. The electrolyte layer enhances thestability of the interference layer of the present invention byprotecting and supporting the membrane that makes up the interferencelayer. Furthermore, the electrolyte layer assists in stabilizingoperation of the device by overcoming electrode start-up problems anddrifting problems caused by inadequate electrolyte. The bufferedelectrolyte solution contained in the electrolyte layer also protectsagainst pH-mediated damage that may result from the formation of a largepH gradient between the hydrophobic interference layer and the electrode(or electrodes) due to the electrochemical activity of the electrode.

Preferably, the coating includes a flexible, water-swellable,substantially solid gel-like film having a “dry film” thickness of about2.5 μm to about 12.5 μm, preferably about 6.0 μm. “Dry film” thicknessrefers to the thickness of a cured film cast from a coating formulationonto the surface of the membrane by standard coating techniques. Thecoating formulation includes a premix of film-forming polymers and acrosslinking agent and is curable upon the application of moderate heat.

Suitable coatings are formed of a curable copolymer of a urethanepolymer and a hydrophilic film-forming polymer. Particularly preferredcoatings are formed of a polyurethane polymer having anionic carboxylatefunctional groups and non-ionic hydrophilic polyether segments, which iscrosslinked in the present of polyvinylpyrrolidone and cured at amoderate temperature of about 50° C.

Particularly suitable for this purpose are aqueous dispersions of fullyreacted colloidal polyurethane polymers having cross-linkable carboxylfunctionality (e.g., BAYBOND®; Mobay Corporation, Pittsburgh, Pa.).These polymers are supplied in dispersion grades having apolycarbonate-polyurethane backbone containing carboxylate groupsidentified as XW-121 and XW-123; and a polyester-polyurethane backbonecontaining carboxylate groups, identified as XW-110-2. Particularlypreferred is BAYBOND® 123, an aqueous anionic dispersion of an aliphatepolycarbonate urethane polymer, sold as a 35 weight percent solution inwater and co-solvent N-methyl-2-pyrrolidone.

Polyvinylpyrrolidone is also particularly preferred as a hydrophilicwater-soluble polymer and is available commercially in a range ofviscosity grades and average molecular weights ranging from about 18,000to about 500,000, under the PVP K® homopolymer series by BASF Wyandotte(Parsippany, N.J.) and by GAF Corporation (New York, N.Y.). Particularlypreferred is the homopolymer having an average molecular weight of about360,000, identified as PVP-K90 (BASF Wyandotte). Also suitable arehydrophilic, film-forming copolymers of N-vinylpyrrolidone, such as acopolymer of N-vinylpyrrolidone and vinyl acetate, a copolymer ofN-vinylpyrrolidone, ethylmethacrylate and methacrylic acid monomers, andthe like.

The polyurethane polymer is crosslinked in the presence of thepolyvinylpyrrolidone by preparing a premix of the polymers and adding across-linking agent just prior to the production of the membrane.Suitable cross-linking agents can be carbodiimides, epoxides andmelamine/formaldehyde resins. Carbodiimide is preferred, and a preferredcarbodiimide crosslinker is UCARLNK® XL-25 (Union Carbide, Chicago,Ill.).

The flexibility and hardness of the coating can be varied as desired byvarying the dry weight solids of the components in the coatingformulation. The term “dry weight solids” refers to the dry weightpercent based on the total coating composition after the time thecrosslinker is included. A preferred useful coating formulation cancontain about 6 to about 20 dry weight percent, preferably about 8 dryweight percent, of polyvinylpyrrolidone; about 3 to about 10 dry weightpercent, preferably about 5 dry weight percent of cross-linking agent;and about 70 to about 91 weight percent, preferably about 87 weightpercent of a polyurethane polymer, preferably apolycarbonate-polyurethane polymer. The reaction product of such acoating formulation is referred to herein as a water-swellablecross-linked matrix of polyurethane and PVP.

D. The Electrolyte Phase

The electrolyte phase is a free-fluid phase including a solutioncontaining at least one compound, usually a soluble chloride salt thatconducts electric current. The electrolyte phase flows over theelectrodes (see FIG. 1C) and is in contact with the electrolyte layer ofthe enzyme membrane. The devices of the present invention contemplatethe use of any suitable electrolyte solution, including standard,commercially available solutions.

Generally speaking, the electrolyte phase should have the same or lessosmotic pressure than the sample being analyzed. In preferredembodiments of the present invention, the electrolyte phase includessaline.

E. The Electrode

The electrode assembly of this invention may also be used in the mannercommonly employed in the making of amperometric measurements. Theinterstitial fluids containing the analyte to be measured is in contactwith a reference electrode, e.g., silver/silver-chloride, and the anodeand cathode of this invention, which are preferably formed of platinum.In the preferred embodiment, the electrodes are connected to a circuitboard in the body of the sensor, the current is read and the informationis radiotransmitted to a receiver. The invention is not limited to thispreferred embodiment. Indeed the membranes of the present inventioncould be used with any form of implantable sensor and adapted to theparticular features of the sensor by one skilled in the art.

The ability of the present device electrode assembly to accuratelymeasure the concentration of substances such as glucose over a broadrange of concentrations enables the rapid and accurate determination ofthe concentration of those substances. That information can be employedin the study and control of metabolic disorders including diabetes.

IV. SENSOR IMPLANTATION AND RADIOTELEMETRIC OUTPUT

Long-term sensor performance is best achieved, and transcutaneousbacterial infection is eliminated, with implanted devices capable ofradiotelemetric output. The present invention contemplates the use ofradiotelemetry to provide data regarding blood glucose levels, trends,and the like. The term “radiotelemetry” refers to the transmission byradio waves of the data recorded by the implanted device to an ex vivorecording station (e.g., a computer), where the data is recorded and, ifdesired, further processed.

Although totally implanted glucose sensors of three month lifetime, withradiotelemetric output, have been tested in animal models at intravenoussites [see, e.g. Armour et al., Diabetes, 39:1519-1526 (1990)],subcutaneous implantation is the preferred mode of implantation [see,e.g., Gilligan et al., Diabetes Care 17:882-887 (1994)]. Thesubcutaneous site has the advantage of lowering the risk forthrombophlebitis with hematogenous spread of infection and also lowersthe risk of venous thrombosis with pulmonary embolism. In addition,subcutaneous placement is technically easier and more cost-effectivethan intravenous placement, as it may be performed under localanesthesia by a non-surgeon health care provider in an outpatientsetting.

Preferably, the radiotelemetry devices contemplated for use inconjunction with the present invention possess features including smallpackage size, adequate battery life, acceptable noise-free transmissionrange, freedom from electrical interference, and easy data collectionand processing. Radiotelemetry provides several advantages, one of themost important of which is the ability of an implanted device to measureanalyte levels in a sealed-off, sterile environment.

The present invention is not limited by the nature of the radiotelemetryequipment or methods for its use. Indeed, commercially availableequipment can be modified for use with the devices of the presentinvention (e.g., devices manufactured by Data Sciences). Similarly,custom-designed radiotelemetry devices like those reported in theliterature can be used in conjunction with the implantableanalyte-measuring devices of the present invention [see, e.g., McKeanand Gough, IEEE Trans. Biomed. Eng. 35:526-532 (1988); Shichiri et al.,Diabetes Care 9:298-301 (1986); and Shults et al., IEEE Trans. Biomed.Eng. 41:937-942 (1994)]. In a preferred embodiment, transmitters areprogrammed with an external magnet to transmit at 0.5 or 5-minuteintervals, depending on the need of the subject; presently, batterylifetimes at transmission intervals of 5 minutes are approximately up to1.5 years.

V. EXPERIMENTAL

The following examples serve to illustrate certain preferred embodimentsand aspects of the present invention and are not to be construed aslimiting the scope thereof.

In the preceding description and the experimental disclosure whichfollows, the following abbreviations apply: Eq and Eqs (equivalents);mEq (milliequivalents); M (molar); mM (millimolar) μM (micromolar); N(Normal); mol (moles); mmol (millimoles); μmol (micromoles); nmol(nanomoles); g (grams); mg (milligrams); μg (micrograms); Kg(kilograms); L (liters); mL (milliliters); dL (deciliters); μL(microliters); cm (centimeters); mm (millimeters); μm (micrometers); nm(nanometers); h and hr (hours); min. (minutes); s and sec. (seconds); °C. (degrees Centigrade); Astor Wax (Titusville, Pa.); BASF WyandotteCorporation (Parsippany, N.J.); Data Sciences, Inc. (St. Paul, Minn.);DuPont (DuPont Co., Wilmington, Del.); Exxon Chemical (Houston, Tex.);GAF Corporation (New York, N.Y.); Markwell Medical (Racine, Wis.);Meadox Medical, Inc. (Oakland, N.J.); Mobay (Mobay Corporation,Pittsburgh, Pa.); Sandoz (East Hanover, N.J.); and Union Carbide (UnionCarbide Corporation; Chicago, Ill.).

Example 1 Preparation of Composite Membrane of the Present Invention

The angiogenic layer may be an ePTFE filtration membrane (Zefluor™, 3.0μm P5PI001, Pall Gelman, Ann Arbor, Mich.) and the bioprotectivemembrane (C30P) may then be coated on the angiogenic layer. For example,the C30P coating solution was prepared by placing approximately 706 gmof dimethylacetamide (DMAC) into a 3 L stainless steel bowl to which apolycarbonateurethane solution (1325 g, Chronoflex AR 25% solids in DMACand 5100 cp) and polyvinylpyrrolidone (125 g, Plasdone K-90D) wereadded. The bowl was then fitted to a planetary mixer with a paddle typeblade and the contents were stirred for 1 hour at room temperature. Thissolution was then coated on the ePTFE filtration membrane by knife-edgedrawn at a gap of 0.006″ and dried at 60° C. for 24 hours.

Alternatively, the C30P solution prepared above can be coated onto a PETrelease liner using a knife over roll coating machine. This material isthen dried at 305° F. for approximately 2 minutes. Next, the Zefluor™ isimmersed in 50:50 (w/v) mixture of tetrahydrofuran/DMAC and then placedupon the coated polyurethane polyvinylpyrrolidone material. Lightpressure atop the assembly intimately embeds the ePTFE into the C30Player. The membrane is then dried at 60° C. for 24 hours.

Example 2 Preparation of the Sensing Membrane

The sensing membrane includes a resistance layer, an enzyme layer, aninterference layer and an electrolyte layer. The resistance layer wasprepared by placing approximately 281 gm of DMAC into a 3 L stainlesssteel bowl to which a solution of polyetherurethaneurea (344 gm ofChronothane H, 29,750 cp at 25% solids in DMAC) was added. To thismixture was added another polyetherurethaneurea (312 gm, Chronothane1020, 6275 cp at 25% solids in DMAC). The bowl was fitted to a planetarymixer with a paddle type blade and the contents were stirred for 30minutes at room temperature. The resistance layer coating solutionproduced is coated onto a PET release liner (Douglas Hansen Co., Inc.Minneapolis, Minn.) using a knife over roll set at a 0.012″ gap. Thisfilm is then dried at 305° F.

The enzyme layer was prepared by placing 304 gm polyurethane latex(Bayhydrol 140AQ, Bayer, Pittsburgh, Pa.) into a 3 L stainless steelbowl to which 51 gm of pyrogen free water and 5.85 gm of glucose oxidase(Sigma type VII from Aspergillus niger) is added. The bowl was thenfitted to a planetary mixer with a whisk type blade and the mixture wasstirred for 15 minutes. Approximately 24 hr prior to coating, a solutionof glutaraldehyde (15.4 ml of a 2.5% solution in pyrogen free water) and14 ml of pyrogen free water was added to the mixture. The solution wasmixed by inverting a capped glass bottle by hand for about 3 minutes atroom temperature. This mixture was then coated over the resistance layerwith a #10 Mayer rod and dried above room temperature preferably atabout 50° C.

The interference layer was prepared by placing 187 gm of tetrahydrofuraninto a 500 ml glass bottle to which an 18.7 gm aliphaticpolyetherurethane (Tecoflex SG-85A, Thermedics Inc., Woburn, Mass.) wasadded. The bottle was placed onto a roller at approximately 3 rpm withinan oven set at 37° C. The mixture was allowed to roll for 24 hr. Thismixture was coated over the dried enzyme layer using a flexible knifeand dried above room temperature, preferably at about 50° C.

The electrolyte layer was prepared by placing 388 gm of polyurethanelatex (Bayhydrol 123, Bayer, Pittsburgh, Pa. in a 3 L stainless steelbowl to which 125 gm of pyrogen free water and 12.5 gmpolyvinylpyrrolidone (Plasdone K-90D) was added. The bowl was thenfitted to a planetary mixer with a paddle type blade and stirred for 1hr at room temperature. Within 30 minutes of coating, approximately 13.1ml of carbodiimide (UCARLNK) was added and the solution was mixed byinverting a capped polyethylene jar by hand for about 3 min at roomtemperature. This mixture was coated over the dried interference layerwith a #10 Mayer rod and dried above room temperature preferably atabout 50° C.

In order to affix this multi-region membrane to a sensor head, it isfirst placed into phosphate buffer (pH 7.4) for about 2 minutes. It isthen stretched over the nonconductive body of sensor head and affixedinto place with an o-ring.

Example 3 In Vivo Evaluation of Glucose Measuring Devices Including theBiointerface Membranes of the Present Invention

In vivo sensor function was determined by correlating the sensor outputto blood glucose values derived from an external blood glucose meter. Wehave found that non-diabetic dogs do not experience rapid blood glucosechanges, even after ingestion of a high sugar meal. Thus, a 10% dextrosesolution was infused into the sensor-implanted dog. A second catheter isplaced in the opposite leg for the purpose of blood collection. Theimplanted sensor was programmed to transmit at 30-second intervals usinga pulsed electromagnet. A dextrose solution was infused at a rate of 9.3ml/minute for the first 25 minutes, 3.5 ml/minute for the next 20minutes, 1.5 ml/minute for the next 20 minutes, and then the infusionpump was powered off. Blood glucose values were measured in duplicateevery five minutes on a blood glucose meter (LXN Inc., San Diego,Calif.) for the duration of the study. A computer collected the sensoroutput. The data was then compiled and graphed in a spreadsheet, timealigned, and time shifted until an optimal R-squared value was achieved.The R-squared value reflects how well the sensor tracks with the bloodglucose values.

To test the importance of the composite membrane of the inventiondescribed in Example 1, implantable glucose sensors including thecomposite and sensing membranes of the present invention were implantedinto dogs in the subcutaneous tissues and monitored for glucose responseon a weekly basis. Control devices including only a bioprotective C30Player (“Control”) were compared with devices including both abioprotective and an angiogenic layer (“Test”), which corresponded tothe composite bioprotective/angiogenic membrane of the device of thepresent invention described in Example 1.

Four devices from each condition were implanted subcutaneously in theventral abdomen of normal dogs. On a weekly basis, the dogs were infusedwith glucose as described above in order to increase their blood glucoselevels from about 120 mg/dl to about 300 mg/dl. Blood glucose valueswere determined with a LXN blood glucose meter at 5-minute intervals.Sensor values were transmitted at 0.5-minute intervals. Regressionanalysis was done between blood glucose values and the nearest sensorvalue within one minute. Devices with an R-squared value greater than0.5 were considered functional. FIG. 3 shows, for each condition, thecumulative number of functional devices over the 12-week period of thestudy. The Test devices performed better than the Control devices overthe entire 12 weeks of the study. All of the test devices werefunctional by week 8. In contrast, none of the control devices werefunctional until week 10, after which 2 were functional for theremaining 2 weeks. The data shows that the use of the inventivebiointerface membrane enables the function of implantable glucosesensors.

The description and experimental materials presented above are intendedto be illustrative of the present invention while not limiting the scopethereof. It will be apparent to those skilled in the art that variationsand modifications can be made without departing from the spirit andscope of the present invention.

What is claimed is:
 1. An implantable sensor for measuring glucose, thesensor comprising: an electrode configured to measure glucose; and amembrane disposed on at least a portion of the electrode, the membranecomprising a first domain that comprises polyurethane andpolyvinylpyrrolidone and a second domain that comprises an enzyme,wherein the first domain is configured to allow transport of glucose andoxygen and prevent entry of inflammatory cells, and wherein the firstdomain is distal to the second domain with respect to the electrode. 2.The sensor of claim 1, wherein the polyurethane ispolycarbonateurethane.
 3. The sensor of claim 1, wherein the firstdomain further comprises at least one material selected from the groupconsisting of polytetrafluoroethylene, polypropylene, polysulfone, andpolyethylene terephthalate.
 4. The sensor of claim 1, wherein the enzymeis glucose oxidase.